Medical radionuclide imaging, commonly referred to as nuclear medicine, is a unique specialty wherein ionizing radiation is used to acquire images which show the function and anatomy of organs, bones or tissues of the body. The technique of acquiring nuclear medicine images entails first introducing biologically appropriate radiopharmaceuticals into the body—typically by injection, inhalation, or ingestion. These radiopharmaceuticals are attracted to specific organs, bones or tissues of interest (These exemplary organs, bones, or tissues are also more generally referred to herein using the term “objects”.). Upon arriving at their specified area of interest, the radiopharmaceuticals produce gamma photon emissions which emanate from the body and are then captured by a scintillation crystal. The interaction of the gamma photons with the scintillation crystal produces flashes of light which are referred to as “events.” Events are detected by an array of photo detectors (such as photomultiplier tubes) and their spatial locations or positions are then calculated and stored. In this way, an image of the organ or tissue under study is created from detection of the distribution of the radioisotopes in the body. Known applications of nuclear medicine include: analysis of kidney function, imaging blood-flow and heart function, scanning lungs for respiratory performance, identification of gallbladder blockage, bone evaluation, determining the presence and/or spread of cancer, identification of bowel bleeding, evaluating brain activity, locating the presence of infection, and measuring thyroid function and activity. Hence, accurate detection is vital in such medical applications.
One particular nuclear medicine imaging technique is known as positron emission tomography, or PET. Positron emission tomography is used to produce images for diagnosing the biochemistry or physiology of a specific organ, tumor or other metabolically active site. The measurement of tissue concentration using a positron emitting radionuclide is based on coincidence detection of the two gamma photons arising from a positron annihilation. When a positron is annihilated by an electron, two 511 keV gamma photons are simultaneously produced and travel in approximately opposite directions. Gamma photons produced by an annihilation event can be detected by a pair of oppositely disposed radiation detectors capable of producing a signal in response to the interaction of the gamma photons with a scintillation crystal. Annihilation events are typically identified by a time coincidence between the detection of the two 511 keV gamma photons in the two oppositely disposed detectors; i.e., the gamma photon emissions are detected virtually simultaneously by each detector. When two oppositely disposed gamma photons each strike an oppositely disposed detector to produce a time coincidence event, they also identify a line(s)-of-response (LOR) along which the annihilation event has occurred.
After being sorted into parallel projections, the LOR defined by the coincidence events are used to reconstruct a three-dimensional distribution of the positron-emitting radionuclide within the patient. In two-dimensional PET, each 2D transverse section or “slice” of the radionuclide distribution is reconstructed independently of adjacent sections. In fully three-dimensional PET, the data are sorted into sets of LOR, where each set is parallel to a particular detector angle, and therefore represents a two dimensional parallel projection p(s, φ) of the three dimensional radionuclide distribution within the patient—where “s” corresponds to the distance of the LOR from the center of the detector and “φ” corresponds to the angle of the detector plane with respect to the x axis in (x, y) coordinate space (in other words, φ corresponds to a particular LOR direction).
Coincidence events are integrated or collected for each LOR and stored in a sinogram. In this format, a single fixed point in f(x, y) traces a sinusoid in the sinogram. In each sinogram, there is one row containing the LOR for a particular azimuthal angle φ; each such row corresponds to a one-dimensional parallel projection of the tracer distribution at a different coordinate along the scanner axis. This is shown conceptually in FIG. 1.
An event is registered if both crystals detect an annihilation photon within a coincidence time window τ (e.g., on the order of 4-5 nsec), depending on the timing properties of the scintillator and the field of view (FOV). The FOV is defined as the volume between the detectors; and a pair of detectors is sensitive only to coincidence events occurring in the FOV. Therefore, the need for physical collimation is eliminated and sensitivity is significantly increased. Accurate corrections (for example, attenuation correction) can be made for the self-absorption of photons within the patient so that accurate measurements of tracer concentration can be made.
The number of time coincidences detected per second within a FOV of a detector is the count rate of the detector. The count rate at each of two oppositely disposed detectors, A and B, can be referred to as singles counts or SA and SB, respectively. The time required for a gamma photon to travel from its point of origin to a point of detection is referred to as the time-of-flight (TOF) of the gamma photon. TOF is dependent upon the speed of light c and the distance traveled. A time coincidence or coincidence event is identified if the time difference between the arrivals of signals in a pair of oppositely disposed detectors is within the coincidence time window τ. In conventional PET, the coincidence detection time window τ is wide enough so that an annihilation event occurring anywhere within the object will produce annihilation gamma photons reaching their respective detectors within the coincidence window. Coincidence time windows of 4.5-12 nsec are common for conventional PET, and are largely determined by the time resolution capabilities of the detectors and electronics.
As illustrated in FIG. 2, if an annihilation event occurs at the midpoint of a LOR, the TOF of the gamma photon detected in detector A (TA) is equal to the TOF of the gamma photon detected in detector B (TB) If an annihilation event occurs at a distance Δx from the midpoint of the LOR, the difference between TA and TB is Δt=2Δx/c, where c is the speed of light. If d is the distance between detectors, the TOF difference Δt could take any value from −d/c to +d/c, depending on the location of the annihilation event.
Time-of-flight (TOF) positron emission tomography (PET) (“TOF-PET”) is based on the measurement of the difference Δt between the detection times of the two gamma photons arising from the positron annihilation event. This measurement allows the annihilation event to be localized along the LOR with a resolution of about 75-120 mm FWHM, assuming a time resolution of 500-800 ps (picoseconds). Though less accurate than the spatial resolution of the scanner, this approximate localization is effective in reducing the random coincidence rate and in improving both the stability of the reconstruction and the signal-to-noise ratio (SNR), especially when imaging large objects. Thus, in TOF-PET, the “TOF” coordinate, Δt, is stored together with s and φ.
There are various types of photosensors that are used or have been studied for use in PET scanners: most commercial systems today are based on photomultiplier tubes (PMTs), which detect the light emitted by a scintillator following a gamma event. Some recent systems and scanner prototypes have used avalanche photodiodes (APDs) to detect the light. Silicon photomultipliers (SiPMs) seem to be a promising sensor type for future detector generations, because they combine some advantages of PMTs (high signal gain and high speed) with those of APDs (small form factor and compatibility with magnetic fields). The output of the above-mentioned sensors is normally an analog signal, where each gamma event leads to a complex signal shape. This shape results from the convolution of the time distribution of photon emission from the scintillator with the temporal response characteristic of the photosensor and possible further broadening by the front-end electronics.
Simultaneous emission (EM) and transmission (TX) data acquisition has a long history in PET scanner design. External positron emitter transmission sources, located outside the emission activity distribution, were exploited. The separation of transmission and emission data involves either dedicated TX detectors [1] or tracing the TX source position in sinogram space [2] during acquisition. With the help of a blank sinogram, acquired from the TX source in the absence of an object, the attenuation coefficients distribution (attenuation map or μ-map) can be reconstructed by iterative algorithm, modeling the TX data statistics [3, 4]. The reconstructed attenuation map is then used for attenuation correction and scatter distribution estimation in emission activity reconstruction. In this way, the reconstruction of two distributions is performed sequentially.
The simultaneous emission activity and attenuation map reconstruction in non-TOF PET was a topic of investigation for more than a decade [5, 6]. The main interest was to exclude the transmission sources completely in order to reduce hardware components. Both distributions were assumed to be reconstructed from a single emission data set. This so-called transmissionless (TX-less) problem has many solutions, even though it was shown that a significant amount of information about attenuation is contained in emission data. The artifact of cross-talking between activity and the attenuation images, when the activity image features propagate to attenuation map images and vice versa, was especially difficult to avoid. Therefore, some constraints on the attenuation distribution, such as partially known distribution and a priori known distribution values, must be used in an attempt to arrive at a proper solution.
The use of acquired TX data in simultaneous EM-TX reconstruction was also considered [7]. Here the main goal was improvement of attenuation map image quality, which suffers from significant noise in TX data alone.
Recent theoretical investigations [8] concluded that both activity and attenuation distributions can be determined from PET TOF data up to the emission image scaling parameter. This encourages TOF TX-less problem investigation in practice [9]. The stability of the solution will need more investigation. In addition to this, the attenuation sinogram cannot be determined outside of the emission sinogram support. Therefore, attenuation map reconstruction still requires a priori knowledge. In a recent TOF practical example [10], spatial attenuation distribution was assumed to be known from a segmented MRI scan. However, the MRI scan alone does not possess information about 511 keV photon attenuation values for a particular object region. Therefore, the MRI image region values were found by TOF simultaneous activity and attenuation reconstruction. This work also showed experimentally that cross talk artifacts, which were one of the sources of solution of non-uniqueness in non-TOF cases, were indeed significantly suppressed.